The present invention relates to the magnetic resonance arts. It finds particular application in conjunction with gradient coils for a magnetic resonance imaging apparatus and will be described with particular reference thereto. However, it is to be appreciated that the present invention will also find application in conjunction with localized magnetic resonance spectroscopy systems and other applications which utilize gradient magnetic fields.
In magnetic resonance imaging, a uniform magnetic field is created through an examination region in which a subject to be examined is disposed. A series of radio frequency pulses and magnetic field gradients are applied to the examination region. Gradient fields are conventionally applied as a series of gradient pulses with pre-selected profiles. The radio frequency pulses excite magnetic resonance and the gradient field pulses phase and frequency encode the induced resonance. In this manner, phase and frequency encoded magnetic resonance signals are generated.
More specifically, the gradient magnetic field pulses are typically applied to select and encode the magnetic resonance with spatial position. In some embodiments, the magnetic field gradients are applied to select a slice or slab to be imaged. Ideally, the phase or frequency encoding uniquely identifies spatial location.
In bore type magnets, linear magnetic field gradients are commonly produced by cylindrical gradient field coils wound on and around a cylindrical former. Discrete coils are wound in a bunched or distributed fashion on a large diameter cylindrical tube, commonly 30 centimeters in diameter or larger.
Historically, gradient coil designs were developed in a "forward approach," whereby a set of initial coil positions were defined and the fields, energy, and inductance calculated. If these quantities were not within the particular design criteria, the coil positions were shifted (statistically or otherwise) and the results re-evaluated. This iterative procedure continued until a suitable design was obtained.
Recently, gradient coils are designed using the "inverse approach," whereby gradient fields are forced to match predetermined values at specified spatial locations inside the imaging volume. Then, a continuous current density is generated which is capable of producing such fields. This approach is adequate for designing non-shielded or actively shielded gradient coil sets.
When designing gradient coils for magnetic resonance imaging, many opposing factors must be considered. Typically, there is a trade off between gradient speed and image quality factors such as volume, uniformity, and linearity. Some magnetic resonance sequences require a gradient coil which emphasizes efficiency, while other sequences are best with a gradient coil which emphasizes image quality factors. For example, a gradient coil which has a large linear imaging volume is advantageous for spine imaging, but is disadvantageous in terms of the dB/dt when switched with a high slew rate.
U.S. Pat. No. 5,736,858 to Katznelson, et al. discloses a magnetic resonance imaging system which has two permanently mounted gradient coil sets where trade offs between linearity and coil performance are taken into account. The gradient coil sets are symmetric coil sets (where the gradient field's sweet spot and the gradient coil's geometric center coincide) which consist of twelve gradient coils (six primary gradient coils and six shield coils). The imaging volume and the performance levels of the two symmetric coils are different. Further, both primary gradient coil sets have different lengths. Specifically, the primary and shield coil combination with better linearity, lower efficiency, and the larger imaging volume is longer than the primary and shield coil combination which has higher efficiency, lower field quality, and a smaller imaging volume. While this design, employing symmetric coils, accounts for some trade offs between efficiency and gradient field quality, it is still problematic.
For very short gradient lengths, the current density of the high efficiency coil increases significantly. The increased current density leads to an increase in the coil's resistance and a shortening of the coil's duty cycle at peak gradient. In this configuration, an extensive cooling scheme, utilizing a multi-layer cooling tube assembly which is difficult to implement, is required. In addition, overlapping of the two symmetric coil sets leads to a significant interaction between them which is measured by the mutual inductance of the coil sets. The mutual inductance of the two coil sets prevents all available power from the amplifier from being directed to the operating modular gradient structure. This reduces the gradient assembly's peak gradient, rise time, and overall slew rate.
The present invention contemplates a new and improved gradient coil set which overcomes the above-referenced problems and others.